About ten years ago, digital X-ray imaging emerged from research units and began to gain ground as a practical imaging method. Ignoring a few exceptions, the earliest digitizing methods were based on a procedure in which X-ray quanta having penetrated the object being imaged were absorbed into a so-called scintillator, which consequently emitted photons, i.e. in a way converted the energy level of the X-ray quanta to a wavelength of light. The photons were subsequently transferred either directly or generally via an optical medium onto a silicon substrate, in which, on being absorbed into the substrate, the photons formed electron-hole pairs, i.e. charges capable of being detected by electrical means. In regard of efficiency, however, such arrangements were relatively modest, and in regard of resolution they were poor because their principle involved the problem of diffusion of photons in the scintillator. On the other hand, as the thickness of the scintillator layer therefore had to be kept rather small, typically at about 10 μm, another consequence was that a large proportion of the X-ray quanta penetrated the scintillator layer and was absorbed into the optical medium. It could even happen that x-ray quanta were absorbed into the silicon substrate itself, which produced a high level of quantum noise in the image information. Often only about 30% of the X-ray quanta could be utilized in image formation.
With the development of the technology, arrangements as described above have yielded a quantum efficiency (dqe) of over 50%. However, the imaging resolution typically achieved by these type of solutions is still only about 10 lp/mm, after which the MTF (Modulation Transfer Function) describing the resolution begins to fall rapidly, being at a level of only about 30% e.g. at frequencies exceeding 5 lp/mm.
In the traditional scintillator material, an X-ray quantum produces, depending on its energy, about 20 photons/keV in the direction of the silicon substrate used as a detector. A proportion of the photons disappears during the passage through the optical fiber transfer medium. Depending on the magnification produced with the fiber optics, typically 10-70% of the photons disappear, which means that even in a favorable case only about 18 photons/keV can be received at the detector. Using a CCD (charge-Coupled Device) sensor having a light quantum efficiency of about 0.3, approximately 6 electrons/keV can be detected. Thus, with X-radiation with an energy level of e.g. 20 keV, the final signal that can be detected is only about 100 electrons/quantum. When such a technique is used, special care has to be taken in designing the sensor electronics to ensure that no more of the information carried by the quanta having penetrated the object gets lost and that every electron detected gets measured.
In the foregoing, reference was made to a CCD sensor, which is a detector generally used in these type of sensor systems. However, a problem essentially associated with CCD technology is the presence of so-called dark current, which arises from the sensor's own surface creep current, which in the course of time leads to an accumulation of signal in the so-called pixel wells of the CCD sensor. Therefore, even if no radiation falls on the sensor, the dark current produces background fog in the image obtained from the sensor, which is a significant disadvantage especially at low signal levels.
Another typical problem with the CCD technique is overexposure. If there is too much signal energy, the charges will start flowing out from the pixel wells into neighboring pixels. This messes up the image and, at the charge transfer phase of the CCD technique, produces so-called blooming.
In all medical X-ray imaging the aim is to keep the dose of radiation the patient is exposed to as low as possible without compromising on image quality. One of the essential factors in this respect is to obtain a quantum efficiency as high as possible, preferably so that all the X-ray quanta having penetrated the target can be made to contribute to image formation. On the other hand, in respect of image quality, it is in many applications important to achieve a maximal imaging resolution. E.g. in mammography, the detection of small micro-calcifications of a size below 100 μm is extremely important.
Technological development has led to a new X-ray quanta detection technique in which the low-efficiency light conversion step is left out altogether. In such a technique, the X-ray quanta are absorbed into a medium (e.g. Ge, Si, Se, GaAs, HgI, CdTe, CdZnTe, PbI) in which they are directly converted into electron-hole pairs. When an intensive electric field is applied across such a medium, the charges produced by the absorption of X-ray quanta can be directed toward pixel electrodes in a way that their lateral movement toward adjacent pixels is in practice prevented. Such a technique provides the advantage of allowing the thickness of the X-ray quanta absorbing layer to be increased up to a theoretical quantum efficiency of 100% without any substantial loss of imaging resolution. To achieve such a level of efficiency, e.g. a silicon layer having a thickness of the order of about 3 mm would be needed, but when e.g. ZdZnTe is used, a thickness of about 0.5 mm will be sufficient for the recovery of all X-ray quanta at 20 keV energy level.
However, the new technology as described above also has certain drawbacks which lead to some limitations regarding the possibilities of utilization of this otherwise excellent technique. As the X-ray quanta absorbing layer in any case has to be relatively thick, it must be arranged in a position as perpendicular as possible to the X-ray beam to ensure that the absorption of quanta at different depths will not lead to their being imaged in the areas of different pixels and therefore to a degradation of lateral imaging resolution. Thus, especially for the purposes of medical X-ray imaging, constructing a so-called full-field sensor using this technology is a somewhat dubious enterprise because the resolution of such a sensor is considerably lower in the edge zones than in the central area. As this problem is common to all sensors based on direct absorption that were known before the present invention, their performance has only been measured in the central part of the image area, where the best imaging result is obtained. By contrast, such a technology would be excellently applicable in so-called narrow-beam scanning imaging, where a narrow sensor can be kept substantially perpendicular to the X-ray beam during the entire scanning process. However, to achieve a sufficient signal sensitivity, scanning imaging would require a sensor structure of a width of several pixels and either the utilization of so-called TDI (Time Delay Integration) technology or a signal reading speed that is unattainable by prior-art solutions in present-day technology.
In full-field sensors like those described above, the image information is generally read using reading electronics provided on the surface of an amorphous silicon substrate. Another possible solution is to use a sensor composed of smaller modules and to implement the reading electronics using CMOS (Complementary Metal-Oxide Semiconductor) technology. In both of these techniques, the image information is generally read either after the imaging and sometimes also during it, pixel by pixel, by addressing one pixel at a time and reading the charge accumulated in it from the edge of the sensor. Especially in large sensors, however, there is the problem of the pixel charge being distributed into reading channel capacitances, producing noise in the signal being measured, and, as stated above, even scanning imaging cannot be implemented because it is not possible to reach the required reading speed by means available today.
When X-ray quanta are converted directly into electron-hole pairs, the signal produced is very large, about 200 keV, as compared with conversion into light. Thus, a 20 keV quantum as mentioned above produces a charge of about 4000 electrons instead of 100 electrons, which means that the problems encountered by the sensor electronics are the converse of the traditional problems. This is to say that the signal now obtained is so large that its processing is becoming difficult. To achieve a sufficient gray scale resolution, an information depth of at least 12-14 bits would be required, which in this case would mean a need to process charges of at least 16-65 Me−. Therefore, the use of e.g. CCD sensors in the reading electronics would be practically impossible because the maximum charge they are able to transfer is only about 500 000 e−. Because of the unsuitability of state-of-the-art reading electronics, this type of detector, though especially suited for scanning imaging, is inapplicable for this purpose as the signal cannot be collected even by the known TDI technique utilizing CCD sensors.
Another disadvantage impeding the use of direct conversion sensors is the limited propagation speed of charge elements in absorbing materials currently used, which gives rise to so-called post-luminescence, which in the worst case may continue for as long as several hours. To deal with this phenomenon, artificial compensation has been employed by taking the information of previous images into account and subtracting it as a function of time from the image taken last. However, this method cannot fully correct the error arising as a result of trapped charges drifting with time even laterally into the area of neighboring pixels.
Digital imaging methods used for medical purposes can be divided into two main categories referred to above, full-field imaging and scanning imaging performed using a narrow sensor. Considering the practical imaging process, full-field imaging corresponds to traditional imaging on a film the size of the entire image area. A distinct drawback associated with this technology is the need for large and therefore very expensive sensors and the need to eliminate the secondary radiation scattering from the object being imaged, requiring the use of complex mechanical grid arrangements. Because of their principle of operation, these grid arrangements also cause a doubling of the dose of radiation needed for imaging.
The narrow sensor used in scanning technology requires some mechanical support, but the costs involved are still considerably lower than those for a full-field sensor. Moreover, scanning imaging requires no grid, so the radiation dose applied to the object to be imaged is correspondingly halved. However, because of the small pixel size (high resolution) needed e.g. in mammography, it would still be necessary to use the TDI method and a sensor having a width of several pixels in order to obtain a sufficient signal with an X-radiation output of a practical magnitude. In state-of-the-art solutions, TDI imaging has generally been implemented using a CCD sensor for signal detection, but in an arrangement based on direct detection such a sensor would not, for the reasons explained above, be capable of reasonable transfer of the signal produced. On the other hand, another state-of-the-art method would be to read the signals detected by pixels connected to an X-Y matrix one at a time by turns, but in the light of the scanning speed and resolution involved in the present applications, this would require 12-bit A/D conversions and recording to be performed at a speed of about 1 ns, which is beyond the capabilities of the technology available today.